Cardiac monitoring to which the present invention relates includes both the determination of heart rate (HR) from electrocardiogram (EKG) signals and the determination of heart stroke volume (SV) from thoracic impedance signals, from which cardiac output (CO) can be estimated.
Heart rate can be determined in a number of ways. The phonocardiogram is considered among the most accurate methods of determining heart rate. However, due to the practical difficulties in using it, the phonocardiogram method is generally not employed for any continuous or long-term monitoring.
Heart rate is most typically determined by the electrocardiogram (EKG). The analog EKG signal typically displays electrocardio events as perturbations referred to as waves. The heartbeat is most clearly reflected in the EKG signal as an R wave peak between a pair of adjoining Q and S wave valleys. The basic and commonly used method of automatically identifying the QRS wave pulses in point is the threshold method in which the rate of voltage change between consecutive data points of the EKG signal is monitored and compared with a threshold value. Slopes exceeding the threshold value are deemed to be associated with a portion of the QRS pulse. While this method commonly detects the interval between consecutive R waves successfully more than eighty percent of the time, it typically has difficulty in dealing with sources of irregular signal components such as pacemakers, muscle noise, 60 Hz interference as well as nearby T or P waves which may also be associated with significant slope changes.
Hemodynamic monitoring of the heart can provide very valuable physiological information regarding the functional state of the myocardium, which is intimately related to its mechanical behavior. The quantitative measurement of blood flow, or the cardiac output (CO), is one of the most useful parameters in assessing cardiac capability, but it is also one of the most difficult to measure. It cannot be accomplished with the electrocardiogram (EKG) which does not reflect the real pumping action of the heart.
Both invasive and non-invasive methods are available for measurement of cardiac output. The invasive methods are considered the most accurate. The risks associated with them are often an unacceptable trade-off, for they require direct access to the arterial circulation. In addition, invasive methods are not suitable for repetitive measurements and normally cannot be performed outside a hospital. Furthermore, invasive methods are very demanding in terms of time consumption and the need for skilled personnel.
Impedance cardiography has been found to be one non-invasive method with the potential for monitoring the mechanical activity of the heart with virtually no risk. It can be conveniently handled by nursing and non-technical staff. It can usually tolerate moderate patient movement and can be left unattended for continuous and long-term monitoring.
U.S. Pat. No. 3,340,867, now U.S. Pat. No. Re. 30,101, to Kubicek et al. discloses an impedance plethysmograph which employs four electrodes, two around the neck and two around the torso of a patient, to provide an impedance difference signal from the two center electrodes. The outermost pair of electrodes apply a small magnitude, high frequency alternating current to the patient while the inner pair of electrodes were used to sense voltage levels on the patient above and below the patient's heart. The impedances of the patient at each of the inner pair of electrodes could be determined from the sensed voltages and known applied currents.
According to Kubicek et al., stroke volume (SV) is related to impedance Z as follows: EQU SV=R(L/Z.sub.o).sup.2 (VET)(dZ/dt.sub.min)
where R is blood resistivity, L is the distance between the inner voltage sensing electrodes, Z.sub.o is the mean thoracic impedance determined from the inner, voltage sensing electrodes, VET is the ventricular ejection time and dz/dt.sub.min is the maximum negative slope change of time-differentiated impedance signal, which is the time-differentiated difference between the impedances determined at the center two electrodes. The above equation is referred to as Kubicek's equation. Cardiac output per minute is stroke volume times heart rate in beats per minute.
The Kubicek equation is based upon a parallel column model of the thorax in which it is assumed:
(1) the thorax is a cylinder, consisting of two electrically conducting tissue paths, also of cylindrical form with uniform cross-sectional areas and homogenous conducting materials, the first path being the blood with a relatively low resistivity and the second path being all other tissues with relatively high resistivities; PA1 (2) the relationship between the maximum impedance change and the stroke volume during the cardiac cycle is linear; PA1 (3) impedance measurements of the individual specific tissue volumes are not very useful in developing the model (the parallel columns model relies on the intact thoracic measurements); and PA1 (4) at 100 kHz frequency, a physiologically safe frequency, the relative thoracic impedance changes are at a maximum, the effects of polarization are negligible, and the reactive component of impedance is small, especially when compared to the real component, so that the reactance could be ignored in determining impedance with only a small error. PA1 (1) poor correlation of the methods and models with the results of the more accepted invasive techniques; PA1 (2) still a relatively high dependance on skilled technical operators; and PA1 (3) still a discomfort to and/or disturbance of patients associated with the use and application of band electrodes.
Yet another model and equation for relating impedance and stroke volume has been proposed by Sramek. According to Sramek, stroke volume (SV) is related to impedance Z as follows: EQU SV=[(O.17H).sup.3 /4.2]*[VET]*[dZ/dt.sub.min /Z.sub.o ]
where H is the patient's height. The Sramek equation is based upon a frustoconical model of the thorax. The Sramek model illustrates some improvement and accuracy over the Kubicek model but its major assumptions are still similar to those of the Kubicek model.
Despite its advantages, impedance cardiography has not been well accepted by clinicians for three primary reasons:
It is believed that poor correlation, the primary reason, can be traced back to a single source, namely the continuing inability to relate impedance cardiography and its mathematical model directly to cardiac physiology.
The following are definitions and abbreviations of some of the terms used frequently herein:
Heart Rate (HR): the number of times the heart contracts each minute.
Ventricular Ejection Time (VET): the time interval of the opening and closing of aortic value during which there is movement of blood out of a ventricle.
Stroke Volume (SV): the volume of blood pumped out by a ventricle (in particular the left ventricle) with each contraction of the heart.
Cardiac Output (Co): the amount of blood pumped out of the heart into the systemic circulation each minute.
Ejection Fraction (EF): the ratio SV/EDV, which is the percentage of blood in a ventricle ejected with each contraction; it is directly related to the strength of the heart with &lt;50% considered abnormal.
End Diastolic Volume (EDV): the volume of blood that fills the ventricle before ejection.
It would be desirable to determine heart rate more accurately than can be determined using the cardiogram threshold method currently employed.
It further would be desirable to provide noninvasive, cardiographic impedance monitoring to estimate stroke volumes, cardiac outputs and related cardiac function parameters which correlate more closely with the stroke volumes, cardiac outputs and the like determined by means of recognized, accepted invasive procedures, but which does not require of operators the technical skills required by current impedance cardiograph systems, and does minimize discomfort to the patient on which the system is used, thereby permitting relatively long-term monitoring of the patient's condition.